The sense of hearing in human beings involves the use of hair cells in the cochlea that convert or transduce acoustic signals into auditory nerve impulses. Hearing loss, which may be due to many different causes, is generally of two types: conductive and sensorineural. Conductive hearing loss occurs when the normal mechanical pathways for sound to reach the hair cells in the cochlea are impeded. These sound pathways may be impeded, for example, by damage to the auditory ossicles. Conductive hearing loss may often be overcome through the use of conventional hearing aids that amplify sound so that acoustic signals can reach the hair cells within the cochlea. Some types of conductive hearing loss may also be treated by surgical procedures.
Sensorineural hearing loss, on the other hand, is caused by the absence or destruction of the hair cells in the cochlea which are needed to transduce acoustic signals into auditory nerve impulses. People who suffer from sensorineural hearing loss may be unable to derive significant benefit from conventional hearing aid systems, no matter how intense the acoustic stimulus is. This is because the mechanism for transducing sound energy into auditory nerve impulses has been damaged. Thus, in the absence of properly functioning hair cells, auditory nerve impulses cannot be generated directly from sounds.
To overcome sensorineural hearing loss, numerous auditory prosthesis systems (e.g., cochlear implant (CI) systems) have been developed. Auditory prosthesis systems bypass the hair cells in the cochlea by presenting electrical stimulation directly to the auditory nerve fibers. Direct stimulation of the auditory nerve fibers leads to the perception of sound in the brain and at least partial restoration of hearing function.
To facilitate direct stimulation of the auditory nerve fibers, a lead having an array of electrodes disposed thereon may be implanted in the cochlea of a patient. The electrodes form a number of stimulation channels through which electrical stimulation pulses may be applied directly to auditory nerves within the cochlea. An audio signal may then be presented to the patient by translating the audio signal into a number of electrical stimulation pulses and applying the stimulation pulses directly to the auditory nerve within the cochlea via one or more of the electrodes.
Current designs of cochlear implant electrode arrays distribute their electrode contacts to various intra-cochlear positions. There is variability in the medial-lateral position, rotational orientation with respect to the modiolus and distribution along the length of the scala tympani. Attempts to control medial-lateral position use preformed designs that aim for a medial position but have no control over length, or rotational orientation in a helical-spiral space. Mid-scala arrays largely control for rotational insertion depth but lead to variability in the medial-lateral direction and also suffer from rotational variability. There are some attempts at accounting for cochlear size through offering a range of array lengths. However, these are only offered for straight arrays that leave the contacts in a lateral position. A design that used a wedge shaped positioner to accommodate a range of cochleae was withdrawn due to safety issues.
A cochlear implant electrode array must distribute its contacts so that discrete bundles of VIIIth nerve fibers are addressed by each contact, hence creating independent channels of information. This is best done by placing the contacts as close to the neural population as possible, meaning adjacent to the medial wall of the cochlea, immediately behind which the target spiral ganglion cells are located. However, the cochlea is buried in the hardest bone of the body, and so offers very limited access.
Also, there is a considerable variation in the size of individual cochleae and in their shape. Usually human cochleae have around 2.5 turns and take a spiral-helical form. To minimize damage to the highly delicate cochlear structures, electrode arrays tend to be introduced via one of the existing openings of the cochlea, the round window, or via a surgically extended round window. This limits the size of the array to under 1 mm diameter. Moreover, the scala tymapani varies in its lateral dimension from under 0.5 mm to over 3 mm along its length, compounded by variation in all other aspects of the cochlea
U.S. Pat. No. 6,266,568 B1 relates to a cochlear electrode array comprising a flexible body on which electrode contacts are carried along a medial side; the flexible body includes an inflatable compartment at the distal end of the electrode array on a side of the flexible body that is opposite the electrode contacts. The electrode array is inserted into the cochlea to a desired depth while the inflatable compartment remains in a deflated state, whereafter a desired modiolus-hugging position is achieved by inflating the inflatable compartment by injecting therein a bio-compatible fluid.
U.S. Pat. No. 7,194,314 B1 relates to an implantable cochlear electrode array comprising a membrane which may be inflated to anchor the array in a position in the cochlea with the electrode contacts pressed into contact with the modular wall, allowing the membrane to seal with the surrounding tissue of the cochlea; the inflatable membrane is inflated once the electrode array has reached its final position.
U.S. Pat. No. 7,822,482 B1 relates to an implantable electrical lead including a rounded array of electrodes which is partly inflatable in order to get closer to a target stimulation site.
U.S. Pat. No. 5,578,084 relates to a cochlear implant electrode array comprising a layer having a controlled rate of expansion when exposed to the water contained in body fluids, making the layer to expand in use and curving the implanted electrode array in order to enable more effective stimulation.